The increasing need for rapid and portable biosensor technology is evidenced by the growing worldwide markets for environmental field testing (FT) (GERSHON J. SHUGAR ET AL., ENVIRONMENTAL FIELD TESTING AND ANALYSIS READY REFERENCE HANDBOOK (2001)) and point-of-care (POC) biomedical diagnostics markets (Kalorama Information, “World Markets for Point of Care Diagnostics,” 13 (2009)), the latter with obvious applications in home health testing (e.g. cholesterol, pregnancy), drugs of abuse screening (e.g. sporting venues, clinic, sobriety check points), and pathogen surveillance for military, homeland security and public health testing. A successful FT/POC platform technology will, in addition to speed, sensitivity and accuracy, be inexpensive and not require extensive training or sophisticated instrumentation for readout. It would utilize a signal transduction strategy that readily extends to the detection of a wide range of targets for which the concentration level that triggers a positive response is tunable for a screening assay and has a wide dynamic range and high target specificity for a quantitative assay. Immunochromatographic test strips comprise many of the commercially-available rapid diagnostics. Signal transduction is based on lateral flow technology that couples a target antibody to a colorimetric agent such as gold nanoparticles which are drawn over capture and control zones by capillary action. Lateral flow devices produce signals detectable by eye but suffer sensitivity and reliability issues and studies indicate that, while simple and rapid, they produce less than acceptable results for wide clinical acceptance (Ferris & Martin, J. Fam. Pract. 34:593-97 (1992); Hook et al., JAMA 272:867-70 (1994); Kluytmans et al., J. Clin. Microbiol. 31:3204-10 (1993)).
In recent years innovation in FT/POC technology development has focused on label free sensors exploiting the unique optical and electrical properties of nanomaterials (Wang et al., Mater. Today 8(5):20-31 (2005); Jain, Clin. Chim. Acta 358(1/2):37-54 (2005)). One such material, electrochemically synthesized porous silicon (PSi) (Bonanno & DeLouise, Anal. Chem. 82:714-22 (2010), holds great promise for FT/POC sensor development. PSi is prepared by anodic electrochemical dissolution of a single crystal silicon wafer in an electrolyte containing hydrofluoric acid (HF) (Jane et al., Trends Biotech. 27:230-39 (2009); Sailor, ACS Nano 1(4):248-52 (2007); DeLouise & Miller, Proc. SPIE 5357:111-25 (2004); Vinegoni et al., “Porous Silicon Microcavities,” in 2 SILICON-BASED MATERIALS AND DEVICES 122-88 (Hari Nalwa ed., 2001)). Etch parameters can be tuned to achieve a high degree of control over pore diameter (10-150 nm) and porosity (20-90%) which are essential properties for fabrication of photonic structures for biosensing applications as they dictate the optical and signal transduction properties and device sensitivity.
There are many advantages of PSi technology for FT/POC sensing applications including inexpensive fabrication, precise control of pore morphology (pore diameter and porosity), intrinsic filtering properties (molecular size selection), high surface area (>100 m2/g), versatile surface chemistry, capacity for label-free colorimetric readout, and compatibility with high throughput array and microfluidic technologies (Bonanno & DeLouise, Anal Chem. 82:714-22 (2010); Sailor, ACS Nano 1(4):248-52 (2007); Jane et al., Trends Biotech. 27(4):230-39 (2009); Bonanno & DeLouise, Biosens. Bioelect. 23:444-48 (2007)). Many proof of principle PSi sensors have been demonstrated for detecting proteins (Ouyang et al., Anal. Chem. 79(4):1502-06 (2007); DeLouise & Miller, Mater. Res. Soc. Symp. Proc. 782:A5.3.1 (2004)), oligonucleotides (Rong et al., Biosens. Bioelect. 23(10):1572-76 (2008); Di Francia et al., Biosens. Bioelect. 21(4):661-65 (2005); Steinem et al., Tetrahedron 60:11259-67 (2004)), enzymes (Kilian et al., ACS Nano 1(4):355-61 (2007); DeLouise & Miller, Anal. Chem. 77(10):3222-30 (2005); DeLouise & Miller, Anal. Chem. 77(7):1950-56 (2005); Orosco et al., Adv. Mater. 18(11):1393-96 (2006)), small molecules (Bonanno & DeLouise, Anal. Chem. 82:714-22 (2010); Lin et al., Biosensor Sci. 278(5339):840 (1997)), and gases (Pancheri et al., Sens. Actuators B 89:237 (2003)). However, little effort has focused on the translation of PSi devices for FT/POC clinical use.
The bio sensor signal transduction principle is based on measurement of refractive index (η) changes. The η of a PSi layer depends on porosity which can be varied precisely between η=3.6 (bulk silicon, 0% porosity), to η=1 (air, 100% porosity). Optical devices (mirrors, microcavities, and rugate filters) are fabricated by etching multilayer structures with alternating porosity (Jane et al., Trends Biotech. 27:230-39 (2009)). These structures function as label-free optical sensors by reporting changes in η (porosity) that result when target binds immobilized receptors. Target binding causes a change in porosity and consequently a change in refractive index (η) that is monitored as a shift in the color of reflected light (i.e., wavelength shift, Δλr) from the sensor. The magnitude of Δλr is a function of the thickness (amount) of the bound material and its refractive index. Target binding decreases porosity, which increases 11, causing a red shift in the optical spectrum.
The magnitude of the optical shift has been shown to be a linear function of pore filling (DeLouise & Miller, Proc. SPIE 5357:111-25 (2004)). Wavelength shift sensitivity (WSS) is a figure of merit specific to each sensor and is measured by displacing air in the pores with liquids of varying refractive index. The WSS value is the slope of the plot of wavelength shift magnitude vs. η. For typical sensors, the WSS values range between 200-400 nm/RIU (DeLouise & Miller, Mater. Res. Soc'y Symp. Proc. 782:A5.3.1 (2003); DeLouise & Miller, Anal. Chem. 77(10):3222-30 (2005); DeLouise & Miller, Proc. SPIE 5357:111-25 (2004)), which translates to detecting a ˜10−3 to 10−4 change in refractive index. This enables target detection sensitivity ranging from mg/ml (Bonanno & DeLouise, Biosens. Bioelect. 23:444-48 (2007) to μg/ml (Dancil et al., J. Am. Chem. Soc'y 121:7925-30 (1999)) or pg/mm2 (DeLouise & Miller, Anal. Chem. 77(10):3222-30 (2005); Lin et al., Science 278:840-43 (1997)) or nM (Kilian et al., ACS Nano 1(4):355-61 (2007)) depending upon the receptor/target system and the assay protocol used (Sailor, ACS Nano 1(4):248-52 (2007); Jane et al., Trends Biotech. 27(4):230-39 (2009)). A much higher limit of detection is desired (picomolar or ng/ml) but this has not yet been achieved with PSi technology. Novel optical signal amplification strategies to increase detection sensitivity and to achieve colorimetric read out by eye would be advantageous.
Devising a strategy to achieve these goals must take into consideration the unique characteristics of the PSi transducer. First, because signal transduction occurs within the porous matrix, the sensor architecture and assay protocol must be designed to overcome the effects of pore blocking, steric crowding, and baseline drift. Baseline drift in the PSi optical response can result from either corrosion of the sensor or from nonspecific adsorption of substances present in complex biological samples (Dancil et al., J. Am. Chem. Soc'y 121:7925-30 (1999); Lees et al., Langmuir 19(23):9812-17 (2003); Canham et al., Physica Status Solidi (A) 182:521 (2000)). Methods to prevent baseline drift are well developed and involve passivating the PSi surface with Si—O or Si—C bond formation (thermal oxidation or hydrosilylation) and utilizing appropriate blocking chemistries and washing protocols (Buriak & Allen, J. Am. Chem. Soc'y 120:1339-40 (1998); Kilian et al., Chem. Commun. 14(6):630-40 (2009); Boukherroub et al., J. Electrochem. Soc'y 149:59-63 (2002); Canham et al., Adv. Mater. 11:1505-09 (1999)). Pore blocking and steric crowding effects are also well understood and can be overcome by tuning the pore diameter and optimizing the surface receptor concentration (DeLouise & Miller, Mater. Res. Soc'y Symp. Proc. 782:A5.3.1 (2003); Bonanno & DeLouise, Langmuir 23:5817-23 (2007)). The latter, unfortunately, may limit the ability to take advantage of the enormous internal surface area of PSi to immobilize a high receptor concentration. Strategies to attain optical signal amplification for improving the limit of detection and colorimetric readout by eye are less developed and constitute an active area of research in the PSi sensor field (Kilian et al., ACS Nano 1(4):355-61 (2007); Orosco et al., Adv. Mater. 18(11):1393-96 (2006); Bonanno & DeLouise, Adv. Funct'l Mater. 20(4):573-78 (2010)).
In traditional bioassay design, signal amplification is commonly achieved using fluorescent or enzymatic secondary reporters (ELISA, PCR). The coupling of enzymatic and/or catalytic reactions to biosensor signal generation is a growing trend (Jane et al., Trends Biotech. 27(4):230-39 (2009); Wang & Lin, Trends Analyt. Chem. 27 (7):619-26 (2008); Jensen & Torabi, J. Optical Soc'y Am. B: Opt. Phys. 3(6):857-63 (1986)). While effective, these methods add significant cost and assay time that label-free technologies seek to overcome for POC applications. Sailor and coworkers have recently demonstrated clever extrapolations of enzymatic signal generation to enhance detection sensitivity of proteases in PSi sensors (Orosco et al., Adv. Mater. 18(11):1393-96 (2006)). In this work a protein layer is coated over a PSi sensor. Protease activity was then detected by measuring optical red shifts (increase in refractive index) due to small peptide fragments (˜7 mM) of the digested protein layer infiltrating the PSi pores. This was followed by the work of Gooding and coworkers (Kilian et al., ACS Nano 1(4):355-61 (2007)) who embedded protein within the PSi matrix and optically detected protease activity (37 nM) by monitoring a blue shift (decreases in refractive index) resulting from protein cleavage and peptide diffusion out of the sensor matrix. These approaches are unfortunately limited to detection of a generic class of enzymes. To overcome these limitations, Voelcker and coworkers have pioneered a label-free optical signal amplification strategy based on inducing PSi corrosion (Steinem et al., Tetrahedron 60:11259-67 (2004); Voelcker et al., Chem Bio Chem. 9:1776-86 (2008)). Formation of a duplex during DNA detection was found to trigger oxidative corrosion of the PSi substrate causing an irreversible increase in porosity and pore size and a profound decrease in refractive index (Steinem et al., Tetrahedron 60:11259-67 (2004)). Detection of DNA at 0.1 amol/mm2 was achieved by this method. This serendipitous effect was later rationally extended by systematically identifying a transition metal complex that could catalyze PSi oxidation. A nickel(II)cyclam derivative was developed as a catalyst label and integrated into a detection assay to achieve amplified detection of biomolecules at submicromolar concentrations (Voelcker et al., Chem Bio Chem. 9:1776-86 (2008)). While this approach is still under development, the irreversible oxidative corrosion of the transducer may prove difficult to control and versatility in target has yet to be demonstrated. While constituting significant advancements, the above mentioned amplification strategies do not directly exploit the fact that the PSi is a volume (porosity) sensitive transducer.
Additionally, clinical and POC diagnostic devices require the specific detection of biological and/or chemical targets at low concentration, in an inexpensive, convenient, reliable, and rapid manner. Many innovative approaches have been reported to address this complex problem yet a need still exists for practical technology solutions. Responsive hydrogels that undergo morphological changes resulting from external stimuli have displayed great promise in chemical sensing (Holtz & Asher, Nature 389:829-32 (1997)) and medical diagnostics (Lapeyre et al., Biomacromolecules 7:3356-63 (2006); Kim et al., Angew. Chem. Int'l Ed. 45:1446-49 (2006); Miyata et al., Nature 399:766-69 (1999)) as well as drug delivery (Kiser et al., Nature 394:459-62 (1998)), tissue engineering (Lutolf et al., Proc. Nat'l Acad. Sci. USA 100:5413-18 (2003)), and microfluidic applications (Yu et al., Appl. Phys. Lett. 78:2589-91 (2001)). Variation of polymer composition, structure, and incorporation of specific functional groups have been exploited to develop hydrogels that respond to an array of biochemical targets including antigen (Yu et al., Appl. Phys. Lett. 78:2589-91 (2001)), DNA (Murakami & Maeda, Biomacromolecules 6:2927-29 (2005)), toxins (Frisk et al., Chem. Mater. 19:5842-44 (2007)), drugs (Ehrbar et al., Nat. Mater. 7:800-04 (2008)), and enzymes (Thornton et al., Chem. Commun. (Camb) 47:5913-15 (2005)). Integration of these smart polymers into specifically engineered sensing systems constitutes an active area of research.
Miniaturization of hydrogel dimensions facilitates reduced response times relative to bulk gel kinetics as required particularly for POC diagnostic testing (Lei et al., Langmuir 20:8947-51 (2004)). Notable success in development of smart hydrogel microlenses into multiplexed stimuli-sensor arrays has been achieved with response time of seconds (Kim et al., Biomacromolecules 8:1157-61 (2007); Dong et al., Nature 442:551-54 (2006)). However, reliance on optical instrumentation to monitor the responses from these microscale devices (change in refractive index or lens radius of curvature) is a drawback for POC applications.
A more attractive approach for POC applications is to integrate smart hydrogels with colloidal crystal arrays (Holtz & Asher, Nature 389:829-32 (1997); Lapeyre et al., Biomacromolecules 7:3356-63 (2006)) or photonic bandgap materials (Segal et al., Adv. Funct'l Mater. 17:1153-62 (2007)). These composite materials potentially enable direct optical detection of hydrogel morphological changes with rapid steady state response times of seconds to minutes. Porous silicon (PSi) is a photonic material that is ideally suited for this application due to its inexpensive fabrication, robust optical transduction, and ease in translation for high-throughput analysis (Chan et al., J. Am. Chem. Soc'y 123:11797-98 (2001); Lin et al., Science 278:840-43 (1997); Bonanno & DeLouise, Biosens. Bioelectron. 23:444-48 (2007); Cunin et al., Nat. Mater. 1:39-41 (2002)). The unique capability of the PSi transducer to report refractive index (η) change that occur within the porous matrix can be exploited to detect target molecules binding directly to the PSi surface or optical changes that occur to a target-responsive gel incorporated into the porous matrix. Chemical and biological sensors have been developed to specifically capture target molecules onto the porous surface area to analyze complex samples in high-throughput and multiplexed assays (Chan et al., J. Am. Chem. Soc'y 123:11797-98 (2001); Lin et al., Science 278:840-43 (1997); Bonanno & DeLouise, Biosens. Bioelectron. 23:444-48 (2007); Cunin et al., Nat. Mater. 1:39-41 (2002)). In addition, visual color readout has been achieved in the detection of protease activity (Orosco et al., Adv. Mater. 18:1393-96 (2006); Gao et al., Anal. Chem. 80:1468-73 (2008)). Protease digestion of a protein layer coated on top of a PSi photonic crystal caused cleavage products to infiltrate the pores producing a large η change that was observed by eye as a color change. These studies highlight the potential for developing PSi photonic sensors for POC diagnostic applications. The capability to easily tune the optical spectrum of the PSi-based 1-D photonic crystal during fabrication facilitates a more deterministic color change combination for portable POC sensing applications. For example, design of a green-to-red color change may be more readily interpreted than a sensor that results in a red-to-deeper-red or blue-to-green color change.
Hydrogel-supported PSi sensors have also been investigated (Segal et al., Adv. Funct'l Mater. 17:1153-62 (2007); DeLouise et al., Adv. Mater. 17:2199-203 (2005); Bonanno & DeLouise, Mater. Res. Soc'y Symp. Proc. 1133:AA01-05 (2008); Bonanno & DeLouise, Proc. SPIE 7167:71670F (2009)). Results show that the sensor maintains the capability to detect small changes in η (10−3-10−4) that result from diffusion of small analytes (DeLouise et al., Adv. Mater. 17:2199-203 (2005)). Composite hydrogel-PSi sensors are also able to detect gel structural changes induced in response to stimuli (temperature and pH) (Segal et al., Adv. Funct'l Mater. 17:1153-62 (2007)) or that result from changes in gel composition (Bonanno & DeLouise, Mater. Res. Soc'y Symp. Proc. 1133:AA01-05 (2008); Bonanno & DeLouise, Proc. SPIE 7167:71670F (2009)). However, incorporation of a bio or chemo responsive hydrogel into a photonic PSi sensor with a tunable target response remains to be demonstrated.
The present invention is directed to overcoming these and other deficiencies in the art.